Ultrasonics has been used in the prior art for purposes of medical diagnosis. Specifically, ultrasonic pulses are transmitted into the body and tissue boundaries produce reflection of the pulses. The transit time of a transmitted and reflected pulse can be measured to provide a determination of the depth of such a boundary.
There is a considerable overlap in the diagnostic uses of ultrasound and computed tomography. Each modality produces cross-sectional images of soft tissue with high spatial resolution and excellent contrast sensitivity or tissue differentiation. However, the imaging mechanisms of the two modalities are entirely different. In computed tomography, an image is mathematically reconstructed utilizing a back-projection algorithm to produce a two-dimensional mapping of X-ray attenuation coefficient. The contrast sensitivity of computed tomography displays local changes in X-ray absorption coefficient of 0.5% of the absorption coefficient of water anywhere in the tomographic image. On the other hand, for diagnostic ultrasound, reflected or scattered mechanical energy is utilized to form images directly. Reflections occur due to the changes in acoustic impedance at every tissue interface. Generally speaking, acoustic impedance of material is the product of its density and the speed of the acoustic waves in the material. In soft tissue imaging, the impedance varies over a range of 60 dB. Even the small changes in the impedance parameter which are associated with soft tissue interfaces (as low as 1 part in a million) are easily detected, resulting in excellent contrast sensitivity.
For each of the computed tomography and ultrasound modalities, a feature of prime importance is the ability to detect lesions of varying size and contrast from the background tissue. For both modalities, the capability of displaying low contrast lesions in a tissue background is limited by two intrinsic imaging parameters, namely: the spartial resolution, and the image noise. These two aspects of image quality have been extensively analyzed and are readily predictable for computed tomography; however, there is very little knowledge concerning image quality characteristics for diagnostic ultrasound devices.
For both the computed tomography and diagnostic ultrasound modality, the limiting three-dimensional spatial resolution for high contrast objects can be described by either the point spread function (PSF) or its Fourier transform, the modulation transfer function (MTF). The analogous two-dimensional spatial resolution within a tomographic image of a given thickness is described by the line spread function (LSF) and its modulation transfer function. The spatial resolution for computed tomography scanners varies to a limited extent over the field of view. The spatial resolution of pulse echo diagnostic ultrasound differs in the axial versus the lateral image dimensions. In the axial dimension, the resolution is determined by the pulse length of the propagating ultrasound pulse. In the perpendicular lateral dimension, due to the wave nature of ultrasound radiation, the spatial resolution is diffraction-limited, depending on the ultrasound wavelength and the f number of a focused transducer. Therefore, for fixed focus ultrasound imaging systems, lateral resolution varies throughout the image field of view. In view of the variation of the spatial resolution with position for both computed tomography and ultrasound, measurements of LSF or MTF should be made at many points in the image and the results averaged to provide a two-dimensional description of spatial resolution.
Spatial resolution or LSF for both diagnostic ultrasound and computed tomography scanners has traditionally been measured by scanning high contrast wires or rods in a water medium. Spatial resolution for abdominal computed tomography systems is on the order of 1 mm square. For diagnostic ultrasound abdominal scanners, axial resolution is optimally about 2 mm and lateral resolution varies from 2 mm at the transducer focus to 1 cm near the transducer and in the far field. In the case of diagnostic ultrasound, high contrast spatial resolution can also be measured by imaging wires or rods suspended in an attenuating tissue equivalent material. Such tissue equivalent resolution phantoms are now commercially available, such as the Model 412 Tissue Phantom manufactured and sold by Radiation Measurements, Inc. of Middleton, Wis., and such as the device illustrated and described in U.S. Pat. No. 4,116,040. Measurements indicate that spatial resolution deteriorates significantly in a tissue medium, primarily due to the frequency-dependent attenuation of tissue and phase-aberration effects of intervening tissue.
The ability of a medical imaging modality to detect a low contrast lesion from a tissue background is limited by the noise in the image. For both computed tomography and diagnostic ultrasound, the noise can be described by the standard deviation of the fluctuation in image intensity from the mean background of an image of a standard uniform test object. Each of the described modalities is subject to electronic noise. Computed tomography also suffers from noise generated due to the algorithm in the mathematical image reconstruction; however, the main noise sources for the two imaging modalities are distinctly different. In computed tomography, as in all radiographic imaging, the primary noise source is quantum mottle, or fluctuations in image background directly related to the photon statistics of image formation. The greater the radiation dose, the less the image noise. In diagnostic ultrasound, the primary noise source is not a function of exposure statistics, but rather is due to coherent speckle, a phenomenon common to all coherent imaging (for example, laser optics). In scanning an abdominal organ, large numbers of scatterers are present in the tissue. Interference effect in the echoes from the multiple scatterers cause severe fluctuations in the image background level which obscure important diagnostic signals.
Due to the restrictions of spatial resolution and image noise for diagnostic ultrasound and computed tomography, low contrast detectability of these modalities is limited. The low contrast performance of these systems can be measured directly using suitable phantoms in the form of objects of varying size and contrast embedded in a tissue-equivalent medium. Several "contrast detail" phantoms have been developed and evaluated for computed tomography applications. An extensive investigation of computed tomography contrast-detail-dose interdependency is described in an article by Cohen, et al., entitled "The Use of a Contrast-Detail-Dose Evaluation of Image Quality in a Computed Tomographic Scanner" appearing in the Journal of Computer Assisted Tomography, Volume 3, pages 189-195, 1979. This paper describes the utilization of the partial volume effect in radiography whereby a phantom was developed containing cylindrical objects varying in contrast from 0.2% to 3% over a range of diameters from 16 mm down to 1 mm. Utilizing this phantom, the threshold of perceptibility of patterns of disks was measured utilizing multiple observers for several computed tomography scanners and dose values. The results for computed tomography indicated the contrast-detail-dose relationship could be divided into (1) a high contrast region (10%-100%) wherein the detection capability was strongly dependent upon the system spatial resolution (MTF) and weakly dependent upon noise (dose) and contrast; (2) a transition contrast region (1%-10%) wherein lesion detectability was dependent upon contrast, noise (dose) and MTF; and (3) a low contrast region (0.1%-1%) wherein the detection was strongly dependent upon image noise, that is, dose. Therefore, for low contrast lesions, image noise becomes the limiting characteristic for detection in computed tomography.
In the field of diagnostic ultrasound, only rudimentary efforts have been made to study the detection capability of low contrast targets. Tissue-equivalent phantoms containing simulated cysts are commercially available and another phantom is marketed containing cylindrical objects whose reflectivity varies from background tissue by 1 dB and 10 dB. These phantoms utilize tissue-equivalent materials of water-based gelatins. Oil-based gels have also been utilized to construct an anthropomorphic ultrasound phantom. However, there has been no attempt to include low contrast objects suitable for quantitative measurements of low contrast detectability of an ultrasound scanner. In each of the prior art phantoms, the variation in contrast or reflectivity is obtained by varying the concentration and particle size of scatterers in the gel matrix of the artificial tissue. For the oil-based gel, polyvinylchloride particle sizes ranging from 100 microns to 260 microns, with concentrations of 0.3 particles per cubic millimeter to 2.0 particles per cubic millimeter demonstrated a reflectivity range from -25 dB to +5 dB relative to the reflectivity of liver.
Thus, where prior art test objects and phantoms enable the evaluation of high contrast resolution in water or tissue-equivalent media, the true efficacy of ultrasound scanners depends upon the ability of such scanners to detect low contrast lesions in tissue. Prior art phantoms simply do not have this low contrast measurement capability. Clearly, then, there is a need for a tissue-equivalent ultrasound phantom capable of permitting measurement of the relationship between threshold detection of lesions of varying size versus contrast (reflectivity).